Squid detected NMR and MRI at ultralow fields

ABSTRACT

Nuclear magnetic resonance (NMR) signals are detected in microtesla fields. Prepolarization in millitesla fields is followed by detection with an untuned dc superconducting quantum interference device (SQUID) magnetometer. Because the sensitivity of the SQUID is frequency independent, both signal-to-noise ratio (SNR) and spectral resolution are enhanced by detecting the NMR signal in extremely low magnetic fields, where the NMR lines become very narrow even for grossly inhomogeneous measurement fields. MRI in ultralow magnetic field is based on the NMR at ultralow fields. Gradient magnetic fields are applied, and images are constructed from the detected NMR signals.

RELATED APPLICATIONS

This continuing patent application claims priority of pending U.S.patent application Ser. No. 10/360,823 filed Feb. 6, 2003, which claimspriority of Provisional Application Ser. No. 60/355,577 filed Feb. 6,2002, which are herein incorporated by reference.

GOVERNMENT RIGHTS

The United States Government has rights in this invention pursuant toContract No. DE-AC03-76SF00098 between the United States Department ofEnergy and the University of California.

BACKGROUND OF THE INVENTION

The invention relates generally to nuclear magnetic resonance (NMR) andmagnetic resonance imaging (MRI), and more particularly to NMR and MRIat ultralow magnetic fields.

Nuclear magnetic resonance (NMR) is a technique for obtaininginformation about atoms and the molecules they form. NMR operates onatoms having nuclei in which at least one proton or neutron is unpaired.This imbalance causes these nuclei to spin on an axis like miniaturetops and gives rise to a magnetic moment, i.e. the nuclei behave likemagnets with north and south poles.

When exposed to an external magnetic field, these spinning magnetsattempt to align their axes along the lines of magnetic force. Thealignment is not exact, however, resulting in a wobbly rotation(precession) about the force lines that is unique for each type ofnuclei. If, while exposed to the magnetic field, the nuclei arebombarded with radio (RF) waves, they will absorb and re-emit energy ata specific frequency according to their rate of rotation. Thisresonating frequency therefore becomes a signature signal by which thenuclei can be identified.

When nuclei absorb the energy of an incoming radio wave, they areknocked out of alignment with the external magnetic field lines. As theysubsequently lose this energy, the nuclei come back into alignment. Therate at which resonating nuclei realign themselves with magnetic fieldlines provides detailed information on their position and motion withrespect to neighboring nuclei. This provides a noninvasive technique tostudy the structural, dynamic, and spatial relationships of atoms in asample of molecules.

NMR has two basic subsets—spectroscopy and imaging. In NMR spectroscopy,the frequency of the incoming radio wave is varied, and all of thedifferent frequencies absorbed and emitted by the nuclei are measured toobtain a resonance spectrum. This NMR spectrum reveals the molecularmakeup of the material down to the respective positions and motions ofthe constituent atoms.

In magnetic resonance imaging (MRI), the frequency of the incoming radiowave is kept constant, but the strength of the external magnetic fieldis varied. The resulting signal corresponds to the total number ofspinning nuclei present in any part of the sample, i.e. the atomicdensity of the sample at that point. Information obtained from an arrayof points can be translated by computer into a recognizable image.

Since the invention of MRI in the early 1970s, MRI scanners havesteadily developed towards higher magnetic field strengths. The enhancedsensitivity attainable at high field makes it possible to resolvefeatures at ever shorter length scales, and enables fast imagingexperiments with close to real-time resolution. State-of-the-artclinical scanners operate at a field strength of 1.5 T, corresponding toa proton Larmor frequency of 64 MHz; currently, there is a drive to gainapproval for 4 T imagers for clinical use. A number of facilities aroundthe world now have 7 T scanners for research purposes.

At the same time, the last three decades have seen continued efforttoward the development of systems for MRI in low magnetic fields. Muchof this work has been motivated by considerations of cost: a commercialfull-body imager operating at 1.5 T costs several million dollars, andthe operation of such a machine places considerable demands on theinfrastructure of the hospital or research facility. In addition, due tothe size and complexity of the high-field system, it must necessarilyremain fixed in one location, and the sample or subject must betransported to the system and inserted into the confining bore of thehigh field magnet; in certain cases this is simply not possible. Alow-cost, portable MRI scanner is extremely appealing, as is an open MRIsystem, which would enable acquisition of MRIs at the same time that amedical procedure is performed. Inexpensive, portable imagers wouldenable MRI to address a wide variety of new problems, potentiallytransforming it from a highly specialized clinical and researchtechnique to a much more widespread, flexible tool for rapid patientscreening and general noninvasive imaging. However, any sort of portableor open MRI system would need to operate at relatively low magneticfield strengths.

Moreover, despite the serious disadvantage of reduced sensitivity, theimages acquired in low field should, in principle, be of higher qualitythan those acquired in high magnetic field. An inevitable drawback ofhigh-field imaging is that of susceptibility artifacts. When aheterogeneous sample is placed in a magnetic field, variations inmagnetic susceptibility over the sample volume give rise to spuriousmagnetic field gradients. When these spurious gradients becomecomparable to the gradients that are used for encoding, the image isseverely distorted. In medical imaging, the presence of dental fillingsor jewelry is enough to destroy the MRI; abrupt changes insusceptibility at solid-liquid and solid-air interfaces inside the body,such as in the sinuses, produce distortions which are more subtle, butwhich nevertheless place strict limits on the achievable spatialresolution. Since the strength of the spurious gradients scales linearlywith the strength of the applied field, it is possible to eliminatesusceptibility-induced distortions entirely by imaging in low magneticfield.

Finally, T₁ contrast in tissue is enhanced in low magnetic field.Because of this, low-field images allow sharper differentiation ofdifferent organs and tissue types, and potentially contain richerinformation than the corresponding images acquired in high field. (It isinteresting to note that, in the early days of MRI, many researcherswere skeptical that high-field MRI would ever amount to a usefulclinical tool, precisely because of the degradation of tissue contrastin high field.)

There have been a number approaches to low-field MRI in recent years.These have generally relied on Faraday detection in a static field oforder 10 mT to 100 mT, which is generated by an electromagnet. The mainobstacle in these studies is the low sensitivity intrinsic to the lowfield experiment. In a different approach, H. C. Seton et al., “A 4.2 Kreceiver coil and SQUID amplifier used to improve the SNR of low-fieldmagnetic resonance images of the human arm,” Meas. Sci. Technol. 8,198-207 (1997) employed a tuned SQUID magnetometer for NMR detection;the SQUID provided an SNR enhancement of a factor of 2.8-4.5 overconventional detection in images acquired from room temperature samplesin a field of 10 mT. In the low-field imaging work of A. Macovski etal., “Novel approaches to low cost MRI,” Magn. Reson. Med. 30, 221-230(1993) and W. Shao et al., “Low readout field magnetic resonance imagingof hyperpolarized xenon and water in a single system,” Appl. Phys. Lett.80, 2032-2034 (2002), spins were prepolarized in a field of 0.3 T, whilethe NMR signals were detected in a much lower field of 30 mT. Here, thehomogeneity of the polarizing field was not crucial, and theprepolarization step led to an enhancement of sample magnetization by anorder of magnitude. Using similar techniques, J. Stepi{haeck over(s)}nik et al., “NMR imaging in the Earth's magnetic field,” Magn.Reson. Med. 15, 386-391 (1990) and G. Planinsic et al., “Relaxation-timemeasurement and imaging in the Earth's magnetic field,” J. Magn. Reson.Ser. A 110, 170-174 (1994), acquired MRIs in the magnetic field of theEarth (B_(Earth)˜50 μT), demonstrating the enhanced T₁ contrastattainable in low-field. In both the works of Macovski et al. andStepi{haeck over (s)}nik et al., however, Faraday detection in the fieldof the Earth entailed substantial signal loss.

Superconducting Quantum Interference Devices (SQUIDs) are sensitivedetectors of magnetic fields based on the quantum mechanical Josephsoneffect. SQUIDs are based on superconductors, whose resistance drops tozero when cooled to a critical temperature Tc. A SQUID is formed byseparating its superconducting material with a very thin insulatingbarrier through which electron pairs can tunnel. This combination ofsuperconducting material and insulating barrier forms a Josephsonjunction, i.e. two superconductors joined by a weak link. The SQUIDconsists of a superconducting ring or square interrupted in two spots byJosephson junctions. When sufficient electrical current is applied tothe SQUID, a voltage is generated across its body. In the presence of amagnetic field, this voltage will change as the strength of the fieldchanges. Thus the SQUID turns a change in a magnetic field, which ismore difficult to measure, into a change in voltage, which is very easyto measure.

For application purposes, SQUIDs are almost always coupled to auxiliarycomponents. To form a magnetometer, a SQUID is connected to a fluxtransformer, a device consisting of a relatively large loop ofsuperconducting material and a much smaller multiturn coil. Since thelarge loop picks up a magnetic field over a much greater area, thesensitivity of the SQUID to changes in magnetic field strength isboosted manyfold.

Originally SQUIDs were made with low Tc superconductors, e.g. niobium(Tc=9.5K), which required cooling with liquid helium. More recently,high Tc SQUIDs have been made, using high Tc ceramic oxidesuperconducting materials, e.g. yttrium barium copper oxide (YBCO)materials (Tc=93K), which only require cooling with liquid nitrogen,which is much less expensive and easier to work with than liquid helium.A high Tc low noise SQUID is described in U.S. Pat. No. 6,023,161 issuedFeb. 8, 2000.

Low transition temperature SQUIDs have been used experimentally todetect NMR and nuclear quadrupole resonance (NQR) signals, e.g. Dinh M.Ton That et al., “Direct current superconducting quantum interferencedevice spectrometer for pulsed nuclear magnetic resonance and nuclearquadrupole resonance at frequencies up to 5 MHz,” Rev. Sci. lnstr. 67,2890 (1996). Low Tc SQUIDs have been used to image polarized helium andxenon at relatively low fields, e.g. M. P. Augustine et al., “Low fieldmagnetic resonance images of polarized noble gases obtained with a dcsuperconducting quantum interference device,” Appl. Phys. Lett. 72 (15),1908 (1998). The feasibility of using a high Tc SQUID to detect NMRsignals has been demonstrated, S. Kumar et al., “Nuclear magneticresonance using using a high temperature superconducting quantuminterference device,” Appl. Phys. Lett. 70 (8), 1037 (1997).

SQUIDs were first used in the 1980s to detect NMR signals in lowmagnetic field. However, the majority of SQUID NMR studies have beenperformed on samples in the solid state, at liquid helium (LHe)temperatures. Recently, there has been increased interest in extendingSQUID NMR techniques to samples in the liquid state, and in particularto systems which are biologically relevant. S. Kumar et al., “BroadbandSQUID NMR with room temperature samples,” J. Magn. Reson. B 107, 252(1995) demonstrated NMR spectra from animal tissue measured at roomtemperature. H. C. Seton et al., ibid., used SQUIDs to image roomtemperature samples in a field of 10 mT, and K. Schlenga et al.,“Low-field magnetic resonance imaging with a high-T_(c) dcsuperconducting quantum interference device,” Appl. Phys. Lett. 75,3695-3697 (1999) and U.S. Pat. No. 6,159,444 issued Dec. 12, 2000, useda SQUID magnetometer fabricated from the high transition temperaturesuperconductor YBCO to image thermally polarized proton samples at roomtemperature in a field of 2 mT. Despite these early efforts, however,SQUID NMR studies of liquid samples remain extremely limited in numberand in scope. The central challenge for SQUID NMR studies of liquids isthat of low sensitivity. Thermal polarizations are two orders ofmagnitude lower at 300 K than at 4.2 K. Moreover, in order to cool theSQUID below its superconducting transition temperature, it is necessaryto thermally isolate the detector from the sample; filling factor istherefore quite low.

The NMR effect is produced by a spin magnetic moment on nuclei in asample. A magnetic field causes the spin magnetic moments to precessaround the field at the Larmor frequency ω which is proportional to themagnetic field.

In low field NMR (typically ≦10 mT) the spin precesses atcorrespondingly low frequencies, typically below 500 kHz, around thefield direction. In conventional NMR, in which a resonant circuit isused to detect the precessing magnetization, the induced voltage signalV is proportional to the spin magnetization M and its rate of change(frequency) ω. Since M is also proportional to the frequency ω, V scaleswith ω². As a result it is difficult to detect NMR signals at low fieldswith a conventional Faraday detector. In contrast, SQUIDs can be used tomeasure magnetic flux directly, resulting in much higher signal to noise(S/N) ratio at low frequencies. However, the use of SQUIDs for NMR/MRIhas heretofore been limited, and not used at ultralow magnetic fields oftens or hundreds of microtesla.

SUMMARY OF THE INVENTION

Accordingly it is an object of the invention to provide method andapparatus for nuclear magnetic resonance and magnetic resonance imagingat ultralow magnetic fields.

The invention is a method and apparatus for the detection of nuclearmagnetic resonance (NMR) signals and production of magnetic resonanceimaging (MRI) by obtaining NMR spectra of liquids in microtesla fieldsusing prepolarization in millitesla fields and detection with an untuneddc superconducting quantum interference device (SQUID). Because thesensitivity of the SQUID is frequency independent, both signal-to-noiseratio (SNR) and spectral resolution are enhanced by detecting the NMRsignal in extremely low magnetic fields, where the NMR lines become verynarrow even for grossly inhomogeneous measurement fields.

The invention operates in ultralow magnetic fields, typically about100-150 μT, but even tens of μT and down to about 1 μT or less, fordetection and typically uses a field of about a few or tens of mT forprepolarization. Sample size can be very small but can be large, i.e.parts of the body. The detector is a SQUID magnetometer designed so thatthe SQUID detector can be very close to the sample, which is at roomtemperature.

A cryogenic insert allows small separations between a sample maintainedat room temperature and a SQUID detector operated in a liquid heliumbath so that NMR signals from samples in the liquid state can bemeasured using a low-T_(c) SQUID. Prepolarization is performed by eitherresonant manipulation of spins in a magnetic field on the order of a fewmT or non-resonant spin manipulation with switched static fields. Thefrequency-independent sensitivity of the untuned SQUID magnetometerallows measurement of NMR signals from the sample in extremely lowmagnetic fields, down to the order of 1 μT, where the proton Larmorfrequency is of the order of tens of Hz. At these fields, NMR linewidthsapproach the lifetime limit, even in grossly inhomogeneous measurementfields. The reduction of signal bandwidth through reduction of thestrength of the measurement field enhances both SNR and spectralresolution. While chemical shift information is lost in the low field,scalar couplings, which are field-independent, are preserved. Thesescalar couplings act as signatures of specific covalent bonds. Thus theinvention includes a simple SQUID-based “bond detector”, which yieldsaccurate information about heteronuclear scalar couplings in microteslafields.

Magnetic resonance imaging in ultralow magnetic field is based on theNMR at ultralow fields. By exploiting the frequency-independent responseof the untuned SQUID magnetometer, it is possible to enhance both theSNR and resolution (in this case, spatial resolution) of magneticresonance images by reducing the strength of the measurement field toeliminate inhomogeneous broadening of the NMR lines.

BRIEF DESCRIPTION OF THE DRAWINGS

FIGS. 1A-D illustrate the narrowing of signal bandwidth throughreduction of measurement field strength.

FIGS. 2A, B illustrate the process of MRI in microtesla magnetic fields.

FIG. 3 shows a cryogenic insert for SQUID-detected NMR spectroscopy ofliquids.

FIG. 4 is a block diagram of the low-T_(c) SQUID spectrometer apparatus.

FIG. 5 shows an NMR pulse sequence for the apparatus of FIG. 4.

FIGS. 6A, B illustrate SNR enhancement through reduction of the strengthof the measurement field.

FIG. 7 is the NMR spectrum of 5 ml of 85% phosphoric acid (H₃PO₄)measured in a field of 2.6 μT.

FIGS. 8A, B illustrate the resolution of scalar couplings in microteslafields.

FIG. 9 shows the magnetic field and magnetic field gradient coils forSQUID-detected MRI.

FIGS. 10A, B are pulse sequences used for SQUID-detected MRI.

FIG. 11 illustrates a combined MEG and MRI apparatus.

DETAILED DESCRIPTION OF THE INVENTION

The invention is directed to certain improvements in NMR/MRI asdescribed herein; other aspects of the NMR/MRI systems are conventionaland not described since they are well known in the art.

A. Concept of Microtesla Field NMR

Line broadening due to field inhomogeneity is a major liability inliquid-state NMR. Spectral resolution, and therefore the informationwhich one can extract about the interaction of nuclei with the localelectromagnetic environment, is ultimately determined by the width ofthe NMR lines: it is necessary that the strength of the interactionsexceed the dispersion of Larmor frequencies in the sample. Moreover, fora fixed sample magnetization, the SNR achieved from a single FID or spinecho signal scales inversely with the inhomogeneously broadened width ofthe NMR line. For these reasons, high-resolution liquid state NMRrequires exquisite field homogeneity. In a conventional high-fieldspectrometer, homogeneity is attained by supplementing the magnet withsophisticated and costly shim coils. In a state-of-the-art high-fieldsystem, homogeneity of a few parts per billion is achieved.

An alternate approach is simply to reduce the strength of themeasurement field. For a fixed relative field homogeneity, absolutehomogeneity is enhanced by decreasing the strength of the measurementfield. Of course, the sensitivity of a conventional NMR detector fallsoff rapidly as field strength is reduced. However, this is not the casefor an untuned SQUID magnetometer, which is sensitive to magnetic flux,rather than the rate of change of magnetic flux. In the context of NMR,this means that, for a fixed sample magnetization, the integratedintensity of the NMR signal—the area under the NMR line—is independentof the frequency of the NMR signal. This makes it possible to performthe following unconventional NMR process.

A sample of nuclear spins is polarized in a magnetic field of the orderof 1 mT. In addition to the polarizing field, a much smaller measurementfield is applied in an orthogonal direction. The polarizing field isthen removed nonadiabatically, inducing precession in the much lowermeasurement field. The sample magnetization is fixed by the strength ofthe polarizing field. The width of the NMR line is determined by theabsolute homogeneity of the measurement field. As the strength of themeasurement field is reduced, the NMR line is compressed into a narrowband and the peak height grows, yielding an enhancement of both spectralresolution and SNR, as shown in FIGS. 1A-D.

FIG. 1A shows a time trace of an NMR signal measured in a highinhomogeneous magnetic field. The dispersion of Larmor frequencies leadsto rapid dephasing of the NMR signal. FIG. 1B shows the NMR spectrumcorresponding to FIG. 1A. FIG. 1C shows a time trace of an NMR signalmeasured in a weak magnetic field with the same relative homogeneity asthat in FIG. 1A, using a detector whose sensitivity is independent offrequency. The NMR signal now appears at much lower frequencies. For afixed sample magnetization, however, the amplitude of the NMR signal isunchanged. Moreover, the effective spin-spin relaxation time T₂* is muchlonger, as the absolute homogeneity of the measurement field has beenenhanced by reduction of the measurement field strength. FIG. 1D showsthe NMR spectrum corresponding to FIG. 1C. The integrated intensity ofthe NMR signal is conserved upon reduction of measurement fieldstrength. Therefore, as the NMR signal is compressed into a narrow band,the peak height grows, leading to an enhancement of both SNR andspectral resolution.

B. Concept of Microtesla Field MRI

There is an intimate connection between the inhomogeneously broadenedwidth of the NMR line and the spatial resolution that can be achieved inan MRI process. The important parameter is the NMR linewidth, which isdetermined by the absolute homogeneity of the measurement field, ratherthan the relative homogeneity. Absolute field homogeneity isconveniently enhanced by reducing the strength of the measurement field.In the case of NMR detection with an untuned SQUID magnetometer, thereduction in measurement field entails no signal loss provided that thesample magnetization is fixed, for example by prepolarization in ahigher field. The basic principles for SQUID-detected MRI in microteslafields is then as shown in FIGS. 2A, B. The sample of nuclear spins ispolarized in field of order tens of mT, corresponding to a polarizationof around 10⁻⁷. However, instead of detecting the NMR signal in a highfield, where it is quite challenging to achieve narrow NMR lines, onemeasures the NMR signal in an extremely low magnetic field, where it ispossible to approach the lifetime limit even for grossly inhomogeneousmeasurement fields. Now it is necessary to apply only modest magneticfield gradients to perform the encoding. As a result, the NMR signal isdispersed over only a narrow band. The NMR transients are thereforedetected with a high SNR, and the time required to acquire the image isrelatively short.

FIG. 2A shows an NMR spectra acquired from a sample (at top) consistingof two separate regions containing nuclear spins which is immersed in amagnetic field which is nominally homogeneous, but which involvesspurious gradients which give rise to inhomogeneous broadening of theNMR ines. At high measurement field strengths, the absolute homogeneityof the field is relatively poor, and the NMR line is relatively broad(center). As the strength of the measurement field is reduced, theabsolute homogeneity of the field is enhanced, and the NMR line isnarrowed (bottom). In the case of fixed sample magnetization anddetection with an untuned SQUID magnetometer, the reduction ofmeasurement field strength also leads to an enhancement of SNR.

FIG. 2B is similar to FIG. 2A, but now a magnetic field gradient G_(z)is applied in order to perform a one-dimensional MRI projection (top).At relatively high measurement field strengths, relatively stronggradients are needed in order to clearly resolve the two spatiallydistinct regions of the sample (center). As a result, the NMR signal isdispersed over a large band, and SNR is poor. At low measurement fieldstrengths, inhomogeneous broadening of the NMR lines is largelyeliminated. As a result, only modest magnetic field gradients are neededto resolve the two spatially distinct regions of the sample. Therefore,the NMR signal remains confined to a relatively narrow band (bottom),and SNR is relatively high.

Even in the absence of homogeneous and inhomogeneous line broadening(infinite T₂ and perfect field homogeneity), MRI resolution isultimately limited by spin diffusion which occurs during the encodingintervals. However, the invention is directed at imaging at the mmlength scale; therefore, spin diffusion does not present a problem. Ofcourse the effects of diffusion can also be overcome by limiting thelength of the encoding intervals, and increasing the strength of theapplied gradients.

C. Experimental Apparatus for NMR

1) Cryogenic Insert

FIG. 3 shows a cryogenic insert 10 in a LHe filled dewar 11 forSQUID-detected NMR spectroscopy of liquids. A cell 12 containing theliquid sample was lowered into the tail section 14 of the insert and wasmaintained at room temperature by a resistive heater. The pickup coil 15of the SQUID gradiometer, with inductance L_(p), was wound around thetail section of the insert. Coils 16, 17 to produce the static magneticfield B₀ (B_(p)) and excitation pulses B₁ (B_(m)) were also located inthe helium bath 18. The dewar 11 was lined with a superconducting Pbshield 19 and the dewar was surrounded by a single-layer mu-metal shield20 to attenuate the Earth's magnetic field and external magneticdisturbances.

In an illustrative embodiment, the inner compartment 21 of the insert,into which the sample was lowered, was 22 mm in diameter. Thiscompartment was surrounded by a liquid nitrogen (LN₂) jacket 22 witho.d. 100 mm to reduce the heat load of the insert on the LHe bath. Atthe 100-mm long insert tail section 14, however, the inner compartmentextended directly into the LHe bath. A single continuous vacuum jacket23 isolated the inner compartment of the insert from the LN₂ jacket (andfrom the LHe bath in the tail region), and isolated the LN₂ jacket fromthe LHe bath; the walls of the vacuum space were silvered, with a slitrunning the length of the insert. The separation between the samplespace and the LHe bath was 5 mm. The insert 10 was surrounded by anumber of styrofoam radiation baffles 24 which were covered withaluminum foil; these extended laterally from the body of the insert tothe neck 25 of the LHe dewar, and served to reduce the heat load on thebath due to gaseous convection and direct radiation from the top of thedewar. To minimize the heat load due to thermal conduction of the Pyrexouter walls of the insert, the boiloff from the LHe bath was extractedfrom the brass top plate 26 of the insert, so that the insert was cooledby the evaporating helium gas. When the tail of the insert was notheated, the system consumed roughly 5 L of LHe per day.

2) SQUID Gradiometer and Readout

A dc SQUID was fabricated using an all-liftoff Nb—AlOx-Nb process. TheSQUID parameters were: 2I_(c)˜5 μA, R_(n)/2˜10 Ω, and L_(s)˜350 pH. Thepeak-to-peak modulation of the SQUID was roughly 40 μV when the SQUIDwas operated in a well-shielded environment. An 11-turn Nb input coilwas integrated onto the SQUID washer.

An illustrative SQUID pickup circuit was configured as a first-orderaxial gradiometer with pickup loop diameter 38 mm and a baseline ofroughly 80 mm, and was wound from 3 mil Nb wire on a fiberglass formthat fit around the tail of the cryogenic insert. For NMR measurementsat low frequency, an untuned, or superconducting, input circuit wasrequired. Superconducting contacts were made from the Nb wire pickupcoil to the on-chip integrated Nb input coil.

A block diagram of the low-T_(c) SQUID spectrometer apparatus 30 isshown in FIG. 4. The dc SQUID 31 consists of a superconducting loop 32interrupted by two Josephson junctions 33. When biased with a currentI_(b) slightly above the critical current of the junctions, the SQUIDacts as a flux-to-voltage transducer. To enhance its sensitivity tomagnetic fields, the SQUID is often operated with a superconducting fluxtransformer 34 consisting of a pickup coil 35 tightly coupled to thesample and an input coil 36 tightly coupled to the SQUID loop 32. Theflux transformer 34 operates on the principle of flux conservation in asuperconducting loop, which involves no frequency dependence. Thus, theSQUID magnetometer can detect broadband at arbitrarily low frequencieswith no loss in sensitivity.

The input coil 36 (with inductance L_(i)) of the transformer 34 wasintegrated onto the SQUID chip; the niobium wire pickup coil 35 (withinductance L_(p)) was wound in a gradiometric fashion around the tailsection of a cryogenic insert. A single-layer solenoid 17 (FIG. 3) ofcopper wire wound on the sample cell produced the polarizing field (B₁or B_(p)). A set of coils 16 (FIG. 3) located in the helium bathprovided the measurement field (B₀ or B_(m)). The SQUID was operated ina flux-locked loop with modulation at 2 MHz, and the signal from theSQUID was amplified, integrated, and fed back to the SQUID as a magneticflux. The voltage across the feedback resistor R_(f) was thusproportional to the applied flux. In this way, the SQUID acted as a nulldetector of magnetic flux.

SQUID operating circuitry is known in the art. In the illustrativecircuit of FIG. 4, the output of SQUID 31 is coupled through transformer37 to amplifier 38 whose output is connected through feedback circuitry42 to feedback coil 41. Circuitry 42 includes a lock-in detector 39which receives inputs from amplifier 38 and oscillator 40. The output ofdetector 39 is integrated by integrator 43. The outputs of integrator 43and oscillator 40 are input into the two inputs of amplifier 44 whichalso has resistor R_(f) connected across the two inputs. R_(f) is alsoconnected to feedback coil 41. The output of amplifier 44 is input tocomputer 45 which controls reset circuit 46 which resets integrator 43.

The small-signal bandwidth of the 2 MHz flux-locked loop was around 700kHz, and the slew rate greater than 10⁶ Φ₀/s. During spin manipulations,the feedback loop was disabled by shorting out the capacitor across theintegrator. The signal from the flux-locked loop passed through asample-and-hold stage (to remove the arbitrary dc level at the loopoutput) and a set of analog filters before digitization. Signalaveraging was performed in software.

3) Static Field and Excitation Coils

For NMR experiments involving resonant spin manipulation, the staticZeeman field was provided by a pair of coils located in the LHe bath andoriented orthogonally to the detection direction. These coils eachconsisted of 67 turns of Cu-clad NbTi wire wound on a 90 mm diameterframe; the separation of the coils was 55 mm. These coils providedroughly 1.2 mT per applied Ampere of current. Resonant pulses wereprovided by a pair of coils oriented along the detection direction andplaced symmetrically with respect to the gradiometric pickup loops ofthe detector, in order to minimize the response of the SQUID to theexcitation. Each coil consisted of 25 turns of insulated Nb wire woundon the 38-mm diameter frame on which the pickup coil was mounted. Theexcitation coils produced a field of roughly 830 μT per applied Ampere.

For NMR experiments involving non-resonant spin manipulation, the67-turn coils located in the helium bath were used to provide ameasurement field of the order of a few microtesla. The spins wereprepolarized along the detection direction in a field of a fewmillitesla; the polarizing field was provided by a one- or two-layersolenoid wound directly on the sample cell from Cu wire.

D. Microtesla Field NMR: Experimental

In the experiments, the polarizing field was applied along the detectiondirection using a one- or two-layer solenoid of copper wire wounddirectly on the sample cell. Currents of order 1 A were used to generatefields of order 1 mT. The measurement field was supplied by the 67-turncoils located in the LHe bath. The sudden switching of the polarizingcoil induced magnetic transients that saturated the detector, giving adeadtime on the order of tens of milliseconds. Spin dephasing duringthis time resulted in signal loss that could be quite significant athigher measurement fields. A spin echo was used to refocus the samplemagnetization. The echo was formed by reversing the direction of themeasurement field, and thus the sense of precession of the nuclearspins.

A pulse sequence for microtesla field NMR is shown in FIG. 5. Thepolarizing field B_(p) of order 1 mT is applied for a time that is longcompared to the spin-lattice time (T₁) of the sample. Followingnonadiabatic removal of the polarizing field, the spins precess in themeasurement field B_(m). As the spins precess, they lose phase coherencedue to inhomogeneity of the measurement field. At a time τ following theremoval of the polarizing field, the direction of the measurement fieldis inverted. Neglecting the effects of diffusion and in the absence ofany background magnetic fields (fields generated by sources other thanthe current in the measurement coil), at each location in the sample,the spins see equal and opposite local magnetic fields before andfollowing the inversion of the measurement field. Therefore the phaseaccumulated by each spin in the interval from t=0 to t=τ is cancelled bythe phase accumulated in the interval from t=τ to t=2τ. At the timet=2τ, phase coherence of the spins is restored, and the echo amplitudeis maximum.

The polarizing field B_(p) is applied along the detection direction,using a solenoid wound directly on the sample cell. The much weakermeasurement field B_(m) is applied in an orthogonal direction. The spinsare polarizedd for a time that is long compared to the spin-latticerelaxation time T₁. Precession is initiated by nonadiabatic turnoff ofthe polarizing field. A spin echo is formed by reversing the directionof the measurement field, and therefore the sense of precession of thenuclear spins.

FIGS. 6A, B demonstrate the SNR enhancement achieved by reduction of themeasurement field strength. FIG. 6A is an NMR (proton) spectrum of 5 mlof mineral oil acquired in a static magnetic field of 1.8 mT withhomogeneity of roughly 10,000 ppm using a conventional Hahn spin echosequence (π/2-τ-π-τ-acq) involving resonant spin manipulation. FIG. 6Bshows the NMR signal from the same volume of mineral oil, measured in afield of 1.8 μT using the sequence of FIG. 5. The sample was polarizedin a field of around 2 mT; the measurement field was applied with thesame magnet used to acquire the spectrum in FIG. 6A. In this case, theproton resonance appears at 77 Hz. The sample magnetizations are thesame in these two experiments; moreover, as the detector is untuned, theareas under the NMR lines are also the same. However, when the protonresonance was lowered from 77 kHz to 77 Hz, the NMR linewidth wascompressed by a factor of 1000, and the peak height grew by the samefactor. Reduction in the measurement field by a factor of 1000 thereforeyielded an SNR enhancement of roughly 1000—note that FIG. 6A representsthe average of 10,000 transients, while FIG. 6B was obtained from theaverage of 100 transients. In measurement fields of the order of 1 μT,SNR was a few tens without signal averaging from samples with a volumeof a few milliliters and a polarization of order 10⁻⁸.

While there is no advantage in either resolution or SNR to be gained inreducing the measurement field past the point where the contribution offield inhomogeneity to the NMR linewidth becomes comparable to thenatural linewidth, proton NMR signals were measured at frequencies aslow as 24 Hz.

E. Multinuclear Studies

When the SQUID magnetometer is operated with an untuned input circuit,it detects broadband. Moreover, as the pulse sequence of FIG. 5 involvesswitched static fields rather than resonant spin manipulation,excitation occurs over a broad band. The technique is therefore ideallysuited to studies of systems containing nuclei with differentmagnetogyric ratios, resonating at different frequencies. FIG. 7 showssimultaneous SQUID detection of ¹H and ³¹P resonances in a field of 2.6μT.

FIG. 7 is an NMR spectrum of 5 ml of 85% phosphoric acid (H₃PO₄)measured in a field of 2.6 μT. The spectrum is the average of 1000transients. The magnetogyric ratios of the spin-1/2 nuclei ¹H and ³¹Pdiffer by a factor of 2.5. The proton resonance appears at 110 Hz; the³¹P resonance is clearly resolved at 44 Hz. The relative intensity ofthe two lines is determined by the different spin densities of the twonuclear species, as well as by the difference in thermal magnetizationsbrought about as a result of the difference in magnetogyric ratios.

F. Scalar Couplings

While all chemical shift information is lost in low magnetic field,scalar (or “J”) couplings, which are field independent, are preserved.These scalar couplings act as signatures of specific covalent bonds. Theenhanced resolution achieved by moving the resonance to low field makesit possible to accurately determine scalar coupling strengths, even ininhomogeneous measurement fields.

FIG. 8A shows,the NMR spectrum obtained from a mixture of methanol andphosphoric acid, measured in a field of 4.8 μT. The proton spectrumconsists of a sharp singlet at 205 Hz. However, when the methanol andphosphoric acid are allowed to react to form the ester trimethylphosphate, scalar coupling to the ³¹P nucleus causes the protonresonance to split into a doublet, with a coupling strengthJ₃[P,H]=10.4±0.6 Hz which is characteristic of this particularnext-next-nearest neighbor interaction (FIG. 8B). The proton doublet iseasily resolved in a field of 4.8 μT, despite a relative fieldhomogeneity of roughly 10,000 ppm over the sample volume.

FIG. 8A is an NMR spectrum of 5 ml of 3 parts methanol, 1 partphosphoric acid (85% in water) measured in a field of 4.8 μT. Thespectrum is the average of 100 transients. Rapid spin exchange with theprotons in water obscures the proton-phosphorous scalar coupling inphosphoric acid, and the proton spectrum consists of a sharp singlet.FIG. 8B is an NMR spectrum of 3 ml of neat trimethyl phosphate(Sigma-Aldrich) measured in a field of 4.8 μT. The spectrum is theaverage of 100 transients. Electron-mediated scalar coupling of the nineequivalent protons to the ³¹P nucleus splits the proton resonance into adoublet, with a splitting that is determined by the coupling strength J.For this particular coupling via three covalent bonds, J₃[P,H]=10.4 ±0.6Hz. Scalar coupling to the nine equivalent protons splits the ³¹Presonance into ten lines; these are below the noise level.

Because electron-mediated scalar couplings between nuclear spins act assignatures of specific covalent bonds, these techniques could form thebasis of a simple low-field NMR “bond detector”, insensitive to chemicalshifts, but yielding accurate information about scalar couplings. Such adetector could be applied to the study of analytes, chemical reactions,and-molecular conformations. For example, the dispersion of J-values forsp³ ¹H—¹³C bonds is approximately 10 times greater than the NMRlinewidths achieved in our experiments. If the values of J-couplings areknown, then pure J-spectra could allow one to assign a number ofmolecular groups. Considering the highly developed techniques forisotopic labeling in biomolecular NMR, the use of such a method forfollowing a “spy nucleus” through bond formation is an appealingpossibility. An illustrative proton spectrum was obtained from 5 ml ofcarbonyl labeled glycine in D₂O in a field of 3.7 μT. The resonance ofthe two equivalent α-protons is split into a doublet due to scalarcoupling to the ¹³C nucleus. The coupling strength was determined fromthe lineshape of the doublet to be J₂[C,H]=5±1 Hz.

G. Experimental Apparatus for MRI: Magnetic Field and Gradient Coils

To perform MRI experiments in extremely low magnetic field, a convenientstarting point was a zero magnetic field region. A set of threeorthogonal cancellation coils 51 (FIG. 9) were used to zero out thefield of the Earth over the measurement region. These coils were woundaround the perimeters of the six faces of a cube 50 that measuredroughly 2 m on a side. Braces integrated into the cube structuresupported a coil assembly 52 which consisted of: (1) a Helmholtz pair53, used to produce a measurement field in the range of μT to tens ofμT; (2) a Maxwell pair 54, used to generate the diagonal componentG_(z)≡dB_(z)/dz of the magnetic field gradient tensor; and two sets ofsaddle coils 55, 56 wound in the Golay geometry, used to generate theoff-diagonal components G_(x)≡dB_(z)/dx and G_(y)≡dB_(z)/dy of thegradient tensor. In addition, the braces supported the fiberglass LHedewar which housed the SQUID sensors. All support structures and coilforms were made from wood, which in many ways was an ideal materialsince wood is non-magnetic and non-conducting. The dimensions of thesystem were selected considering the eventual imaging of human subjects:an average-sized adult can fit (not too uncomfortably) into themeasurement region at the center of the cube.

A schematic of the coil system 52 is shown in FIG. 9. Following theconvention in the magnetic resonance community, the z axis lies alongthe measurement field direction; the x axis is chosen to coincide withthe vertical direction (the detection direction).

FIG. 9 shows the magnetic field and magnetic field gradient coils forSQUID-detected MRI. Six 100-turn coils 51 wound on each face of a cube,2 m on a side, were used to cancel the magnetic field of the Earth. Themeasurement field B₀ was produced by a Helmholtz pair 53 arranged in thecenter of the cube. The diagonal component G_(z) of the first-ordermagnetic field gradient tensor was produced by a Maxwell pair 54; theoff-diagonal gradients G_(x) and G_(y) were produced by saddle coils 55,56 wound in the Golay geometry (for each set of Golay coils, only two ofthe four saddle coils are shown). The detector was a second-order axialSQUID gradiometer oriented in the vertical (x) direction, and which washoused in a LHe dewar suspended in the center of the cube.

1. Cancellation Coils

Each of the six cancellation coils consisted of 100 turns of 18 gaugecopper wire wound in a groove cut along the outer edge of a wooden framewhich made up one of the six faces of the cube. The coils on oppositefaces of the cube were wired in series. Each pair had a resistance ofroughly 30 Ω, and generated a magnetic field change of around 50 μT perapplied Ampere, so that only fractions of Amperes and modest voltageswere required to cancel the three components of Earth's field. Heatingin the cancellation coils was negligible. (For the record, a total ofabout 2.7 miles of 18 gauge Cu wire—approximately 60 lbs—went into-thecancellation coils).

2. Measurement Field and Gradient Coils

The measurement field was produced by a Helmholtz pair with radius 0.6 marranged in the center of the cube. Each coil consisted of 20 turns of18 gauge copper wire. The diagonal component G_(z) of the gradienttensor was generated by a Maxwell pair with radius 0.6 m mounted outsidethe measurement field coils. Each coil consisted of 20 turns of 18 gaugewire. The off-diagonal gradients G_(x) and G_(y) were generated bysaddle coils wound in the Golay geometry from 20 turns of 22 gauge wire;the radius of curvature of the Golay coils was 0.6 m. These coils werewound in grooves cut along the outer edges of circular pieces ofplywood. Per Ampere of applied current, these coils generated: 1) B₀=30μT; 2) G_(z)=50 μT/m; and 3) G_(x)=G_(y)=50 μT/m.

3. Polarizing Coil

In order to enhance the sample magnetization, the spins wereprepolarized in a magnetic field of the order of tens of millitesla.Precession was induced either by non-adiabatic turnoff of the polarizingfield (in which case the spins were polarized in a direction orthogonalto the measurement field), or by adiabatic turnoff of the polarizingfield followed by resonant excitation in the much lower measurementfield (in which case the spins were typically polarized along themeasurement field direction). In the first instance, the polarizingfield had to be switched in a time which was short compared to theLarmor period in the measurement field; in the second instance, theswitching requirements were less strict: turnoff of the polarizing fieldneeded to occur simply in a time which was short compared to thespin-lattice relaxation time T₁. In either case, the requirement ofrapid field switching, coupled with practical limits to the level ofJoule heating which could be tolerated in the magnet wire, governed thedesign of the polarizing coil.

For rapid switching it is advantageous to use relatively heavy gaugewire and a relatively small number of turns. The polarizing magnetconsisted of two poles wound from 18 gauge copper wire on fiberglassframes. On each pole, the magnet windings filled a 17 mm wide channel,with a 15 mm radius for the inner windings and a 70 mm radius for theouter windings. The separation of the two magnet poles was roughly 100mm. A total of 510 turns was used for each magnet pole. The field at thecenter of the coils was approximately 4 mT per applied Ampere; the totalresistance of the magnet pair was approximately 6 Ω, while theinductance was roughly 100 mH. The heavy formvar insulation of themagnet wire seemed to tolerate Joule heating at a level of 100 W permagnet pole. To achieve higher polarizing fields, however, it would benecessary to water cool the magnet. While the homogeneity of thepolarizing field was quite poor, this of course had no effect on NMRlinewidth or SNR.

The condition for nonadiabatic turnoff of the polarizing field can bewritten as follows:${{\frac{\mathbb{d}B}{\mathbb{d}t} ⪢ {B\quad\omega}} = {\gamma\quad B^{2}}},$where B is the instantaneous strength of the magnetic field. Formeasurement fields in the microtesla range, it is easy to meet thenonadiabatic switching condition, as the term on the right hand sidebecomes extremely small as the strength of the measurement field isreduced. Indeed, in our experiments, it was sufficient to drive thepolarizing coil with a shaped voltage pulse from a Techron 7700amplifier. At the turnoff of the pulse, the coil-would discharge with anL/R time governed by the self-inductance of the coil (again, around 100mH) and the output impedance of the amplifier (a few tens of ohms).Driving the polarizing coil in this “controlled voltage” mode had theadvantage that, following the turnoff of the polarizing field, amechanical relay in series with the polarizing coil could be opened todecouple the SQUID gradiometer from noise generated by the Techronamplifier, as well as from the thermal noise of the polarizing coilitself. With this simple switching scheme, the nonadiabatic criterionwas satisfied for proton Larmor frequencies extending to around 2 kHz.As the strength of the measurement field was increased beyond thisvalue, some loss in signal occurred, due presumably to adiabaticreorientation of the sample magnetization during turnoff of thepolarizing field. At higher measurement frequencies, a scheme involvingadiabatic removal of the polarizing field and resonant spin excitationwas used.

4. Spin Echo Coil

Resonant spin echo pulses were produced with a pair of coils orientedalong the y direction. The coils each consisted of 10 turns of copperwire, wound on circular fiberglass forms with radius 60 mm which weremounted rigidly to the polarizing coil forms. In the case of the NMR andMRI experiments performed in this system, resonant echoes werepreferable to echoes formed by inverting the direction of themeasurement field, as the latter type of echo does not refocus dephasingfrom external magnetic field gradients (either the gradients G_(x),G_(y), G_(z) which were used for the encoding, or spurious gradients dueto nearby magnetic objects). For Larmor frequencies of the order of 1kHz, a typical π pulse consisted of two or three cycles, and requiredcurrents of order 10 mA.

H. Experimental Apparatus for MRI: SQUID Sensors

The SQUID system which was housed in the LHe dewar in the center of thecube consisted of four channels: a single large-area sensing channel andthree orthogonal magnetometer references. The pickup coil of the sensorwas configured as a second-order axial gradiometer oriented along the xaxis (vertical direction); the gradiometer was therefore sensitive tothe component d²B_(x)/dx² of the second-order magnetic field gradienttensor. The pickup loops of the gradiometer were wound from 3 mil Nbwire in grooves which were carefully machined in a cylindrical G10fiberglass form. The radius of each pickup loop was roughly 15 mm, andthe gradiometer baseline was 50+50 mm. The form on which the pickup loopwas wound fit snugly into the tail section of the fiberglass LHe dewar.The distance from the sensing loop of the gradiometer to the sample,which was positioned directly under the tail of the LHe dewar, wasroughly 10 mm. The sensing SQUID was housed in a superconducting boxmachined from a solid block of Pb and located in the LHe dewar at adistance of approximately 300 mm from the sensing coil of thegradiometer. Superconducting contact from the Nb wire pickup coil to the11-turn Nb input coil integrated on the SQUID chip was achieved usingthe solder blob technique. The sensing SQUID was operated in aflux-locked loop with modulation at 2 MHz.

The NMR detector was a second-order axial gradiometer with an overallbaseline of 100 mm. The probe also incorporated a three-axis referencemagnetometer; the reference SQUIDs were mounted on orthogonal faces of aG10 fiberglass cube which was positioned inside the cylindrical form onwhich the pickup coil of the gradiometer was wound.

The inductance of the-sensing SQUID was roughly 350 pH. This-yields forthe 11-turn input coil a self inductance L_(i)=40 nH and a mutualinductance to the SQUID of M=3.9 nH. For a 1+2+1 turn second-ordergradiometer with pickup loop radius r=15 mm, L_(p)=700 nH. The aboveparameters yield a gradiometer sensing area A_(sense) of approximately3.7 mm². Note that the sensing area of the gradiometer could be improvedsubstantially by proper matching of L_(i) and L_(p). In particular, byincreasing the number of turns in the input coil to around 45, so thatL_(i)˜700 nH, the sensing area of the gradiometer could be increased toroughly 8 mm². However, even for a reduced sensing area of 3.7 mm²,system noise was dominated by external sources of thermal noise andinterference, not by the intrinsic noise of the detector. In this case,inductance mismatch does not degrade SNR, as signal and noise areattenuated equally.

To measure the balance of the gradiometer with respect to uniform fieldsin the three orthogonal directions, known, highly uniform fields wereapplied to the sensor using the cancellation coils wound on the faces ofthe cube, while monitoring the flux that was coupled to the sensingSQUID. The typical gradiometer balance was one part in a few hundred forboth in-plane and out-of-plane fields. To achieve this level of balance,however, parasitic inductance associated with the Nb wire leads whichran between the gradiometer loops, and which were not twisted together,was minimized.

Finally, the probe incorporated a three axis SQUID magnetometerconsisting of three Nb—AlOx-Nb SQUIDs mounted on three orthogonal facesof a G10 fiberglass cube. This reference cube was mounted inside thefiberglass form on which the gradiometric pickup coil of the sensingSQUID was wound. Each of the reference SQUIDs had an effective area ofroughly 0.03 mm², and was operated in a flux-locked loop with modulationat 100 kHz. Analog subtraction of the reference signals from thegradiometer signal was used to improve the balance of the gradiometer,in an attempt to further reduce the contribution of distant noisesources to the system noise.

I Choice of Measurement Field

In all of the NMR and MRI experiments performed in this system, thesample magnetization was enhanced by prepolarization in a field of theorder of tens of millitesla. In the case of fixed sample magnetizationand detection with an untuned SQUID magnetometer or gradiometer, theintegrated intensity of the NMR signal is independent of the strength ofthe measurement field. This allowed considerable freedom in the choiceof measurement field strength for the experiments. The choice ofmeasurement field was dictated by the following considerations: 1)inhomogeneous broadening due to spurious gradients generated bymeasurement field coil; 2) level of environmental interference and noiseover the signal band; and 3) the strength of fields generated bygradient coils, relative to the measurement field strength (possibleimage distortion due to concomitant gradients).

J. NMR Experiments

The procedure for performing NMR experiments in the cube was as follows.First, appropriate currents were passed through the cancellation coilsto zero out the static magnetic field of the Earth over the measurementregion. For this purpose, a three-axis fluxgate magnetometer placeddirectly under the tail of the cryostat was used to monitor the field inthe measurement region. The desired measurement field, of the order ofmicrotesla or tens of microtesla, was then applied with the 1.2 mdiameter Helmholtz pair. After allowing some time for the measurementfield to stabilize, the sample was placed under the tail of the cryostatand tuned the SQUID gradiometer.

As mentioned above, two different polarization and excitation schemeswere employed in the experiments: at lower measurement fields(corresponding to Larmor frequencies below about 2 kHz), spin precessionwas initiated by nonadiabatic turnoff of the polarizing field; at highermeasurement fields, the polarizing field was reduced to zeroadiabatically, and a resonant π/2 pulse was used to induce precession.In both cases, a spin echo, which was formed with a resonant π pulse,was typically detected.

FIGS. 10A, B show pulse sequences used for SQUID-detected MRI. In thesequence of FIG. 10A, precession is induced by the nonadiabatic turnoffof the polarizing field B_(p) (of the order of tens of millitesla),which is oriented in a direction orthogonal to the measurement fieldB_(m). The measurement field is static; a spin echo is formed with aresonant π pulse. In the sequence of FIG. 10B, the strength of thepolarizing field B_(p) is reduced adiabatically to zero, and precessionis induced with a resonant π/2 pulse. In this case, the polarizing fieldis applied along the measurement field direction, to avoid signal lossdue to imperfect fulfillment of the adiabatic switching criterion. Inboth cases, turnoff of the polarizing field is accomplished in a time oforder 10 ms; adiabaticity or nonadibaticity of the turnoff is determinedby the Larmor period of the nuclear spins in the measurement field. Thenonadiabatic turnoff was used for measurement fields below about 50 μT;at higher measurement field strengths, the switching of the polarizingfield was adiabatic, and only the adiabatic sequence was used.

In the experiments, polarizing field strengths were of the order of afew tens of millitesla, and the polarizing field was applied for a timewhich was long compared to the spin-lattice relaxation time T₁ of thesample (polarizing intervals of 2-3 s were used for water, and of around100 ms for mineral oil). Following turnoff of the polarizing field andapplication of the echo pulse, mechanical relays in series with thepolarizing and echo coils were opened. This was done to isolate theSQUID gradiometer from the noise of the amplifiers used to drive thepolarizing and echo coils, as well as to prevent the flow of Nyquistnoise currents in these coils. The flux-locked loop was enabled shortlyafter transmission of the echo pulse. The output of the flux-locked loopwas passed through sample-and-hold and filtering stages and thendigitized; signal averaging was performed in software.

When the external gradients were properly compensated, NMR linewidths oforder 1 Hz were obtained from a test sample of water. With a polarizingfield of roughly 40 mT, an SNR ratio of 50 was obtained in a single shotfrom a sample of around 20 ml of water.

K. MRI Experiments

For imaging experiments, projection reconstruction was chosen overFourier reconstruction due to the limited hardware requirements as wellas the ease with which this encoding scheme is adapted to SQUIDdetection. MRIs were acquired from tap water or mineral oil phantoms.The phantoms were chosen to have translational symmetry in onedirection, which was arranged to coincide with the axis of thesecond-order gradiometer (the x axis). The lateral (y-z) dimensions ofthe phantom were chosen to roughly match the dimensions of the sensingloop of the gradiometer (30 mm diameter). The gradient component G_(x)was nulled out using the appropriate set of Golay coils; appliedgradients G_(y) and G_(z) were used to encode in the y-z plane. Noattempt was made to perform slice selection, although of course thesensitivity of the SQUID gradiometer fell off rapidly as a function ofthe separation of the source from the sensing loop. Therefore the MRIsobtained were two-dimensional cross sectional images of the phantoms.

The first experiments were carried out with tap water phantoms. In thiscase, the long T₂ (of the order of seconds) made it possible to achievevery narrow lines of 1-2 Hz in low field, and therefore to take fulladvantage of signal bandwidth narrowing in order to enhance SNR andresolution. However, the long T₁ necessitated polarization times of theorder of seconds. Due to the long polarizing interval, overall imageacquisition times were rather long, of the order of a few hours.

Subsequent imaging experiments were performed with mineral oil, which inany case is a better match to human tissue in terms of its NMRrelaxation properties. Here T₂ is shorter, and (lifetime limited) protonlinewidths of around 5 Hz were achieved. The broader NMR lines requiredencoding over a broader band in the MRI experiments. On the other hand,the shorter T₁ of mineral oil enabled a significant decrease in thepolarizing interval, to around 100 ms, resulting in a substantialreduction in image acquisition times.

Preliminary experiments such as this demonstrated the soundness of theconcept of microtesla field MRI. However, the acquisition times could besubstantially reduced by reduction of the system noise.

Further experiments have been done with a low-noise G10 fiberglasscryostat that can accommodate a gradiometer with pickup coil diameter upto 160 mm. A second-order gradiometer was constructed with an overallbaseline of 150 mm, and pickup loop diameter 65 mm. The 1+2+1 turnpickup coil has an inductance of roughly 1.7 μH, yielding a gradiometersensing area A_(sense)=7.6 mm². In addition, in order to eliminate thecontribution to the system noise from distant sources of noise andinterference, an eddy current shield surrounds the SQUID MRI system. Theshield consists of a single layer of ⅛″ (roughly 3 mm) 5052 Al plate(ρ=4.8 μΩ cm); its dimensions are 8′×8′×12′. At a frequency of around 5kHz, the shield is roughly two skin depths thick, corresponding to anattenuation of interfering fields in the relevant frequency range by anorder of magnitude. Inside this eddy current shield and with the new,large-area gradiometer, the thermal noise from the CTF dewar was around2.5 fT/Hz^(1/2) at a frequency of 5 kHz; at the moment, the system noiseis dominated entirely by thermal noise from the dewar. The current noiselevel of 2.5 fT/Hz^(1/2) is roughly a factor of 6 lower than the noiselevel achieved in the prior measurements, corresponding to a potentialreduction in MRI acquisition time by a factor of 36.

The larger pickup coil of the new gradiometer will allow acquisition ofMRIs from larger samples, and some in vivo imaging experiments. However,the larger area of the sensing loop entails a significant loss ofsensitivity to dipole sources (and in MRI, one is concerned with dipolesources). The signal flux coupled to a pickup loop of radius r from adipole goes as r⁻¹ (this is seen easily if the problem is viewed interms of mutual inductances). On the other hand, in the case of a systemwhose noise is dominated by the magnetic field environment, the noiseflux coupled to the pickup loop will go as r². Therefore, in more thandoubling the diameter of the gradiometer sensing loop, the sensitivityto magnetic dipoles decreased by more than a factor of 2³=8, roughlycanceling the advantage gained in reducing the magnetic field noise ofthe system. Due to these considerations, it will ultimately be desirableto replace the single large-area sensor with an array of smallersensors, designed for maximum sensitivity to localized sources. This isin fact the current trend in high-field clinical MRI, where system noiseis dominated by inductively coupled losses from the human body.

L. Conclusion

Thus the invention provides method and apparatus for obtaining nuclearmagnetic resonance (NMR) spectra of liquids in fields of a fewmicrotesla using prepolarization in fields of a few millitesla anddetection with a dc superconducting quantum interference device (SQUID).Because the sensitivity of the SQUID is frequency independent, bothsignal-to-noise ratio and spectral resolution are enhanced by detectingthe NMR signal in extremely low magnetic fields, where the NMR linesbecome very narrow even for grossly inhomogeneous measurement fields.

The technique of bandwidth-narrowing through measurement in microteslafields can also be applied to magnetic resonance imaging (MRI). In MRI,the ultimate spatial resolution is determined by the width of the NMRline in the absence of applied magnetic field gradients. For MRIperformed in low field with linewidths approaching the lifetime limit,relatively high spatial resolution is achievable with modest magneticfield gradients. As a result the NMR signal is dispersed over only anarrow band, resulting in a high S/N ratio, and thus a short acquisitiontime. Furthermore, distortions due to spurious gradients generated as aresult of spatial variations in magnetic susceptibility can be minimizedin low field.

The invention thus leads to low-cost, portable MRI scanners as well asopen MRI systems, which would have significant advantages.

The low-field MRI techniques of the invention could readily be adaptedfor use with existing commercial SQUID systems designed to measure weakmagnetic signals generated by electrical currents in the human body.Here, the frequency-independent response of the untuned SQUIDmagnetometer would be exploited to detect biomagnetic signals atfrequencies in the range of Hz to tens of Hz, as well as to perform MRIat proton Larmor frequencies on the order of kHz. It would beparticularly useful to develop a SQUID system which could be used forboth MRI and magnetoencephalography (MEG). In MEG, an array of low-T_(c)SQUID sensors (up to 300 channels in state-of-the-art systems) is usedto measure weak neuromagnetic signals (of the order of tens of fT) inorder to map out electrical activity in the brain; the technique israpidly becoming more widespread in clinical use, especially for studiesof epilepsy and for pre-surgical screening of brain tumors. The abilityto perform MRI would greatly increase the utility of the SQUID system,creating a versatile, new tool for studies of the human brain.

FIG. 11 illustrates a patient under a combined MEG/MRI helmet orapparatus 60.

Changes and modifications in the specifically described embodiments canbe carried out without departing from the scope of the invention whichis intended to be limited only by the scope of the appended claims.

1. A method for nuclear magnetic resonance (NMR) of a sample,comprising: prepolarizing nuclear spins in the sample in a milliteslamagnetic field; detecting nuclear magnetic resonance (NMR) signals fromthe sample in a microtesla magnetic field with an untuned low criticaltemperature (Tc) superconducting quantum interference device (SQUID)magnetometer.
 2. The method of claim 1 wherein the detecting magneticfield is in the range of about 150 μT to about 1 μT.
 3. The method ofclaim 2 wherein the prepolarizing magnetic field is tens of mT.
 4. Themethod of claim 1 wherein the SQUID is maintained substantially atliquid helium temperature and the sample is at room temperature.
 5. Themethod of claim 1 wherein the prepolarization is performed by resonantspin manipulation and the prepolarization and detection fields areorthogonal.
 6. The method of claim 1 wherein the prepolarization isperformed by nonresonant spin manipulation and the prepolarization anddetection fields are parallel.
 7. The method of claim 1 wherein thedetection magnetic field is substantially inhomogeneous.
 8. The methodof claim 1 further comprising operating the SQUID in a flux locked mode.9. The method of claim 8 wherein the output of the SQUID is amplified,integrated, and then fed back to a flux modulating coil positioned nearthe input coil of the SQUID.
 10. The method of claim 1 furthercomprising performing magnetic resonance imaging (MRI) of the sample byforming an image from the detected NMR signals.
 11. The method of claim10 further comprising applying encoding magnetic field gradients to thesample.
 12. The method of claim 1 further comprising simultaneouslydetecting multinuclear species.
 13. The method of claim 1 furthercomprising obtaining scalar (“J”) coupling information from the NMRsignals.
 14. Apparatus for nuclear magnetic resonance (NMR) of a sample,comprising: prepolarization coils for providing a millitesla magneticfield for prepolarizing nuclear spins in the sample; measurement coilsfor providing a microtesla magnetic field for detecting nuclear magneticresonance (NMR) signals from the sample; an untuned low criticaltemperature (Tc) superconducting quantum interference device (SQUID)magnetometer for detecting nuclear magnetic resonance (NMR) signals fromthe sample.
 15. The apparatus of claim 14 further comprising a fluxlocked loop connected to the SQUID.
 16. The apparatus of claim 14wherein the detecting magnetic field is in the range of about 150 μT toabout 1 μT and the prepolarizing magnetic field is tens of mT.
 17. Theapparatus of claim 14 further comprising a liquid helium dewar in whichthe SQUID is mounted to maintain the SQUID at liquid helium temperaturewhile the sample is at room temperature.
 18. The apparatus of claim 14further comprising means to perform magnetic resonance imaging (MRI) ofthe sample by forming an image from the detected NMR signals.
 19. Theapparatus of claim 18 wherein the means to perform MRI include gradientfield coils.
 20. The apparatus of claim 18 further comprising a SQUIDarray for performing magnetoencephalography.